The present invention is directed to a transparent solid scintillator material where the material contains praseodymium as an activator for rapid conversion of exciting radiation, specifically x-rays, to scintillating radiation. The present invention is also directed to an exciting energy detection device, specifically an x-ray detection device, which incorporates the scintillator material.
A luminescent material absorbs exciting energy of one type, and then emits electromagnetic energy. If the exciting energy is electromagnetic radiation, the luminescent material will absorb the exciting electromagnetic energy in one region of the electromagnetic spectrum and generally will emit energy in another region of the electromagnetic spectrum. A luminescent material in powder form is called a phosphor, while a luminescent material in the form of a transparent solid body is called a scintillator.
An impurity activated luminescent material is normally one in which a non-luminescent host material has been modified by including ions of an activator species within the host material. With an impurity activated luminescent material, the host material lattice of the luminescent material absorbs the incident photon and the absorbed energy may be accommodated by the activator ions or it may be transferred by the lattice to the activator ions. The luminescent activator ions are then raised to a more excited state. In returning to their less excited state, the ions emit a photon of luminescent electromagnetic energy.
The exciting energy for a luminescent material may come in the form of electrons, positrons, electromagnetic radiation, or other forms of energy. X-ray scintillators are formed from scintillator material which absorbs x-ray radiation and then emits electromagnetic scintillating radiation. Typically the electromagnetic scintillating radiation is in the visible region of the electromagnetic spectrum. A typical x-ray detector using an x-ray scintillator includes a scintillator material which absorbs x-rays and a photodetector which detects scintillating radiation emitted from the scintillator due to the absorption of the x-rays. In general, if the scintillator material absorbs more x-rays, the scintillator material will emit more scintillating radiation to the photodetector, and the signal output from the photodetector indicative of the scintillating radiation will be greater.
X-ray detectors using scintillator materials are often used in x-ray diagnostic devices such as medical diagnostic equipment or baggage inspection equipment. One particular application of x-ray scintillator materials is in medical imaging detectors such as computed tomography (CT) equipment. In a typical CT scanning system, an x-ray source and an x-ray detector array are positioned on opposite sides of the subject and rotated around the subject in fixed relation to each other. In a CT scanning system using a scintillator material the scintillator material of a cell or element absorbs x-rays incident on that cell and emits light which is collected by a photodetector for that cell. During data collection, each cell or element of the detector array provides an output signal representative of the scintillating radiation intensity in that cell of the array. These output signals are processed to create an image of the subject in a manner which is well known in the CT scanner art.
For medical imaging systems, such as CT scanners, the scintillator material should have a number of important characteristics. First, in x-ray based CT systems, it is desirable to absorb substantially all of the incident x-rays in the scintillator material in order to minimize the x-ray dose to which the patient must be exposed in order to obtain the computed tomography image. In order to collect substantially all of the incident x-rays, the scintillator material should be of a sufficient density to efficiently stop the x-ray photons.
Secondly, the scintillator material should have a good quantum conversion efficiency, i.e., the ratio of the number of scintillating radiation photons emitted to the number of x-ray photons absorbed should be high. If the quantum efficiency of a scintillator material is high, more scintillating radiation photons will be emitted, and consequently the number of scintillating radiation photons to be detected by the scintillating radiation detector will be advantageously higher.
A number of materials are known to be useful as a scintillator material for medical imaging detector applications, for example, yttrium gadolinium oxide (Y,Gd)2O3 doped with europium, and Gd2O2S doped with praseodymium. Both of these materials have the density required to efficiently stop the x-ray photons, as well as efficient conversion of the x-ray energy absorbed in the scintillator material into visible light emission from the Eu and Pr activators. (Y,Gd)2O3:Eu, however, suffers from a long decay time of the Eu emission which limits its usefulness in applications when the x-ray signal is rapidly changing, as might be encountered in fast scan CT systems. Gd2O2S:Pr has a hexagonal crystal structure which generates scattering in a polycrystalline ceramic body thus reducing the efficiency with which the generated light can escape the body and strike the scintillating light photodetector.
Further complicating the choice of a scintillator material appropriate for a fast scan CT systems is the difficulty of predicting a priori which materials will have the properties mentioned above which are important for a fast scan CT system. For example, much characterization of luminescent materials has been done using ultraviolet (UV) radiation as the stimulating radiation because ultraviolet radiation is more easily produced than x-rays and is generally considered less harmful. Unfortunately, there are a number of materials which are luminescent in response to UV radiation stimulation which are not luminescent in response to x-ray stimulation. Consequently, for many materials, even that luminescent data which is available provides no assurance that the material will luminesce in response to x-ray stimulation.
Afterglow is a scintillator material property which is often undesirable in a CT, or other scanning systems. Afterglow in an x-ray detecting scintillator is the phenomena that luminescence from the scintillator due to x-ray excitation can still be observed a long time after the x-ray radiation is absorbed by the scintillator. Upon absorbing x-ray radiation the scintillator will emit light where the intensity of the light decays rapidly at an exponential rate. This first exponential rate of decay is the primary decay rate. Additionally, the scintillator will emit a lower intensity light where the light intensity decays much more slowly than the primary decay rate light. The more slowly decaying light is termed afterglow.
It is desirable that the afterglow be small so that the photodetector of a scintillator system is able to distinguish between the scintillating light of a present x-ray stimulation from that of a prior x-ray stimulation. This is particularly important for fast scanning systems where the time between sequential stimulations is small.
Radiation damage in an x-ray scintillator material is the characteristic of the scintillator material in which the quantity of light emitted by the scintillator material in response to the stimulating x-ray radiation changes after the material has been exposed to a high radiation dose. The radiation damage may be expressed as the percent decrease in intensity of the stimulated scintillating light emitted before and after a high dose of X-ray radiation exposure of the scintillator material. Thus, usefulness of an x-ray scintillator material will be greater for a lower radiation damage value, because the scintillator material can withstand a higher x-ray radiation dose without a change in the proportionality calibration between stimulating x-ray radiation intensity and scintillating light emitted.
In view of the foregoing, it would be desirable to provide a scintillator material which has a fast primary decay time and is highly transparent. It would also be desirable to provide a scintillator material with a small afterglow and low radiation damage.
According to one embodiment of the invention a transparent solid scintillator material is provided. The solid scintillator material has a cubic garnet host. The solid scintillator material also contains praseodymium which acts as an activator.
According to another embodiment of the invention, a computed tomography system is provided. The computed tomography system includes an x-ray source, a transparent solid scintillator material, and a scintillating radiation detector which is optically coupled to the transparent solid scintillator material for detecting scintillating radiation. The solid scintillator material has a cubic garnet host. The solid scintillator material also contains praseodymium which acts as an activator.
According to yet another embodiment of the invention, a fast response x-ray detector is provided. The fast response x-ray detector includes a transparent solid scintillator material and a scintillating radiation detector optically coupled to the transparent solid scintillator material for detecting the scintillating radiation. The solid scintillator material has a cubic garnet host. The solid scintillator material also contains praseodymium which acts as an activator.
According to yet another embodiment of the invention, a process of making a transparent solid polycrystalline ceramic scintillator material is provided. In this process a phosphor powder of a cubic garnet host with praseodymium as an activator is formed. The phosphor powder is pressed to from a powder compact, and then sintered.
According to yet another embodiment of the invention a phosphor powder is provided. The phosphor powder has a cubic garnet host, and contains praseodymium which acts as an activator. The phosphor powder also includes an afterglow reducing element distributed within the cubic garnet host.